WO2019142720A1

WO2019142720A1 - Buckle, state of alertness determination system, and state of alertness determination method

Info

Publication number: WO2019142720A1
Application number: PCT/JP2019/000515
Authority: WO
Inventors: 青木　洋
Original assignee: ＪｏｙｓｏｎＳａｆｅｔｙＳｙｓｔｅｍｓＪａｐａｎ株式会社
Priority date: 2018-01-18
Filing date: 2019-01-10
Publication date: 2019-07-25
Also published as: JP2019122652A; JP6975649B2

Abstract

Provided is a buckle comprising: a sensor for outputting an output signal that varies according to the breathing of a crew member in a vehicle; a detection unit for detecting a frequency component of the breathing from the output signal by means of frequency analysis; and a determination unit for determining the state of alertness of the crew member on the basis of the frequency component detected by the detection unit. Also provided is a state of alertness determination method, in which: a sensor provided in a buckle outputs an output signal that varies according to the breathing of a crew member in a vehicle; a detection unit detects a frequency component of the breathing from the output signal by means of frequency analysis; and a determination unit determines the state of alertness of the crew member on the basis of the frequency component detected by the detection unit.

Description

Buckle, awake state determination system and awake state determination method

The present invention relates to a buckle, an awake state determination system, and an awake state determination method.

BACKGROUND Conventionally, there is known a technique for determining the awakening state of a person based on respiration data of the person acquired by a respiration sensor. For example, a device that determines an awake state using, as an index, an average value of RI (Respiration Interval) indicating a breathing interval or RrMSSD (Respiration root Mean Square Successive Difference) indicating a variation of RI in a predetermined interval of about 120 seconds. (See, for example, Patent Document 1).

International Publication No. 2015/060268

A statistical index in a given section such as an average value of RI or RrMSSD is suitable for representing an average feature of respiration in that section, and is effective for rough determination of arousal state. However, if information on time-series characteristics of respiration in a section (for example, periodic change in increase or decrease of respiration in the section, a pattern of increase or decrease, etc.) is rounded off by statistical processing in that section, determination of arousal state Accuracy may be reduced.

Thus, the present disclosure provides a buckle, an awake state determination system, and an awake state determination method that can determine an awake state with high accuracy.

The present disclosure
A sensor that outputs an output signal that changes in response to the breathing of a vehicle occupant;
A detection unit that detects frequency components of the respiration from the output signal by frequency analysis;
And a determination unit that determines an awake state of the occupant based on the frequency component detected by the detection unit.

Also, the present disclosure
A buckle having a sensor that outputs an output signal that changes in response to the breathing of a vehicle occupant;
A detection unit that detects frequency components of the respiration from the output signal by frequency analysis;
The awake state determination system includes: a determination unit that determines an awake state of the occupant based on the frequency component detected by the detection unit.

Also, the present disclosure
A sensor provided on the buckle outputs an output signal that changes in response to the breathing of the vehicle occupant,
The detection unit detects frequency components of the respiration from the output signal by frequency analysis;
The determination unit provides an awake state determination method that determines the awake state of the occupant based on the frequency component detected by the detection unit.

According to the present disclosure, the awakening state can be determined with high accuracy.

It is a figure showing an example of composition of a seat belt device. It is a block diagram showing an example of composition of a buckle in a 1st embodiment. It is a flowchart which shows an example of the respiration signal extraction process which a detection part implements. It is a flowchart which shows an example of the breathing cycle statistical processing which a detection part implements. It is a flowchart which shows an example of the respiration frequency component ratio detection process which a detection part implements. It is a figure which shows an example of the respiration signal before normalization. It is a figure which shows an example of the respiration signal after normalization. It is a figure for demonstrating some simple methods about the normalization process which carries out amplitude shaping | molding. It is a figure which shows an example of the statistical analysis of the respiratory cycle fluctuation in a predetermined area. It is a wave form diagram showing an example of each signal detected from the driver under driving. It is a figure which shows an example of the relationship between the frequency range of LFR and HFR, and the pass frequency of a low pass filter and a high pass filter. It is a figure which shows an example of the structure which calculates LFR, HFR, and RLHR using a low pass filter and a high pass filter. It is a figure which shows function f1, f2, f3 for the production | generation of a convolution filter. The schematic diagram of F2-F1 filter characteristic and F3 filter characteristic is shown. It is a figure showing superposition with a convolution filter, a low pass filter, and a high pass filter. It is a figure which shows an example of the result of having calculated RLHRn (normalization output of a respiration frequency component ratio) using the filter of FIG. It is a figure which shows an example of the time-sequential data of 20 respiration cycles. It is a figure which shows an example of frequency distribution (occurrence frequency) which makes a horizontal axis a period area of a 0.5 second interval about 20 respiration cycle data. It is a figure which shows an example of frequency distribution (occurrence frequency) which divided the horizontal axis of frequency distribution of FIG. 18 in the frequency, and rearranged it. It is a figure which shows an example of the breathing waveform of repetition cycle 6 second. It is a figure which shows an example of the result of having carried out spectrum analysis of the waveform of FIG. It is a flowchart which shows an example of the simple estimation method of a respiration frequency component ratio. It is a flowchart which shows an example of the simple estimation method of a respiration frequency component ratio. It is a figure which shows an example of the result of having calculated RLHRn (normalized output of a respiration frequency component ratio) using the simple estimation method of a respiration frequency component ratio. It is a block diagram showing an example of composition of a buckle in a 2nd embodiment.

Hereinafter, embodiments according to the present invention will be described with reference to the drawings.

FIG. 1 is a view showing an example of the configuration of a seat belt device. The seat belt device 1 is an example of an in-vehicle system mounted on a vehicle. The seat belt device 1 includes, for example, a seat belt 4, a retractor 3, a shoulder anchor 6, a tongue 7, and a buckle 8.

The seat belt 4 is an example of a seat belt that restrains the occupant 11 sitting on the seat 2 of the vehicle, and is a belt-like member that can be taken up by the retractor 3 so as to be drawn out. The seat belt is also referred to as webbing. The belt anchor 5 at the front end of the seat belt 4 is fixed to the seat 2 or the vehicle body in the vicinity of the seat 2.

The retractor 3 is an example of a winding device that enables the seat belt 4 to be wound or pulled out, and the seat belt 4 is pulled out of the retractor 3 when a deceleration equal to or greater than a predetermined value at the time of a vehicle collision is applied to the vehicle. Limit that. The retractor 3 is fixed to the seat 2 or the vehicle body in the vicinity of the seat 2.

The shoulder anchor 6 is an example of a belt insertion tool through which the seat belt 4 is inserted, and is a member for guiding the seat belt 4 pulled out from the retractor 3 toward the shoulder of the occupant 11. The shoulder anchor 6 is fixed to the seat 2 or the vehicle body in the vicinity of the seat 2.

The tongue 7 is an example of a belt insertion tool through which the seat belt 4 is inserted, and is a component slidably attached to the seat belt 4 guided by the shoulder anchor 6.

The buckle 8 is a component to which the tongue 7 is detachably connected, and is fixed to, for example, the seat 2 or a vehicle body near the seat 2.

The buckle 8 has a main body 8a and a stay 8b. The main body 8a is a portion to which the tongue 7 is detachably connected. The stay 8 b is an example of a support member that supports the main body 8 a of the buckle 8. The stay 8 b is fixed to the seat 2 or the vehicle body in the vicinity of the seat 2.

When the tongue 7 is connected to the buckle 8, a portion of the seat belt 4 between the shoulder anchor 6 and the tongue 7 is a shoulder belt portion 9 that restrains the chest and shoulders of the occupant 11. In a state where the tongue 7 is connected to the buckle 8, a portion of the seat belt 4 between the belt anchor 5 and the tongue 7 is a lap belt portion 10 that restrains the waist of the occupant 11.

FIG. 2 is a block diagram showing an example of the configuration of the buckle 8 in the first embodiment. In the first embodiment, the buckle 8 includes a sensor 20 and an estimation unit 30. The sensor 20 outputs an output signal that changes in accordance with the respiration of the occupant 11 of the vehicle. The estimation unit 30 estimates the awake state of the occupant 11 based on the output signal output from the sensor 20.

The estimation unit 30 includes, for example, at least one computer including at least one CPU (Central Processing Unit) and at least one memory. A specific example of a computer is a microcomputer. Each function of the estimation unit 30 is realized by processing that at least one program causes the CPU to execute. The program is readably stored in the memory. The estimation unit 30 includes a plurality of functional blocks of a detection unit 40, a determination unit 70, and an output unit 80.

The detection unit 40 detects respiration information representing the respiration state of the occupant 11 from the output signal of the sensor 20. The determination unit 70 determines the awake state of the occupant 11 based on the respiration information detected by the detection unit 40. The output unit 80 outputs the determination result of the awake state by the determination unit 70 to the external device of the buckle 8 by wire or wirelessly. The external device executes predetermined control (for example, control to warn an occupant, control to support the traveling of a vehicle, and the like) based on the determination result.

Here, there is a technique for estimating the awakening state of a person using heart rate information representing a heart rate state of the person instead of estimating the respiration state representing the person's respiration state. For example, heart rate variability HRV derived from the heart rate interval RRI may be used to estimate the arousal state of a person. The heart rate variability HRV contains information representing autonomic nerve activity. In particular, LF / HF, which is an indicator of the balance of activities of sympathetic nerve and parasympathetic nerve, is used as an important indicator of nerve activity related to arousal state. The sympathetic nerve has an action of enhancing the activity of the brain and the body. For example, when it gets up in the morning, the sympathetic nerve becomes active and a person becomes awakening state. The parasympathetic nerve has the effect of suppressing the brain's excitement, and when the parasympathetic activity is dominant, the arousal state decreases. For example, when sleeping at night, this state occurs. Therefore, when the value of LF / HF is large, sympathetic nerves are dominant, and when the value of LF / HF is small, parasympathetic nerves are dominant.

This LF / HF index measures the power spectrum of the waveform after frequency analysis by analyzing the frequency of the heartbeat fluctuation HRV representing time-series fluctuation of the heartbeat interval RRI in a predetermined section (for example, 1 to 2 minutes) Calculated by The integrated value of the power spectrum amplitude of low frequency components in the range of 0.04 Hz to 0.15 Hz on the frequency axis of the waveform after frequency analysis is LF, and the power spectrum amplitude of high frequency components in the range of 0.15 Hz to 0.5 Hz The integrated value of is HF. HF is an indicator of parasympathetic activity, and increases as parasympathetic activity becomes active. On the other hand, LF is known as an index indicating the activity of sympathetic nerve and parasympathetic nerve, and LF is increased by active sympathetic activity or active parasympathetic activity. Therefore, by taking the ratio of LF to HF (LF / HF), it is possible to estimate the size of the activity balance between the sympathetic nerve and the parasympathetic nerve, and this LF / HF is also used to estimate the arousal state. For example, if the sympathetic nerve becomes active, LF / HF becomes large, and if the parasympathetic nerve becomes active, LF / HF becomes small. That is, when the LF / HF becomes larger than a predetermined determination threshold because the sympathetic nerve is dominant, it can be determined that the person is in the awake state. The frequency division of LF / HF differs somewhat depending on the document. In the text, figures of 0.04 Hz, 0.15 Hz and 0.5 Hz are used, but the frequency division is not limited to these figures.

On the other hand, human respiration repeats inspiratory and exhalation in a cycle of 2 seconds to 6 seconds to introduce oxygen into the body. Inhalation causes air to be taken into the lungs by dilating the lungs and creating negative pressure therein. Exhalation exhales air by compressing the lungs and providing a positive pressure therein. At this time, the concentration of oxygen taken into the blood from the lungs decreases at negative pressure and increases at positive pressure. Then, depending on the breathing cycle, the concentration of oxygen taken into the blood from the lungs changes up and down.

The heart has a function to send oxygen taken into blood to the whole body. The sinus node responsible for the contraction of the atrium of the heart contracts the atrium as a pacemaker of the heart and sends blood to the ventricle. The atrioventricular node contracts the ventricle at a time delay from the sinus node's excitation and sends blood throughout the body with strong pressure. The sinus node responds to the oxygen concentration in the blood, and has an automatic adjustment function to suppress the heartbeat when the oxygen concentration is increased, if the oxygen concentration is decreased. As described above, since heart rate variability is affected by respiration through changes in oxygen concentration, heart rate variability contains respiratory variation information, and it can be said that heart rate variability and respiration are correlated.

Further, as a method of detecting a heartbeat, there is contact detection using an electrocardiograph, pulse measurement which measures a pulse wave, and the like. There is also a technology for detecting a minute movement of the body surface with a microwave or an electrostatic sensor to detect a heartbeat. However, in the on-vehicle environment, even when performing non-contact detection for detecting the heart rate without making the occupant feel troublesome, unlike the resting state on the bed, there is vibration or movement of the body due to running of the vehicle. It becomes difficult to stably detect minute movements of the

On the other hand, respiration is large on the surface of the body compared to the heartbeat, and the respiration frequency does not overlap so much with the vehicle vibration frequency region of 1 Hz or more (the heartbeat frequency almost overlaps with the vehicle vibration frequency region). Therefore, in the on-vehicle environment, detection of respiratory information representing a respiratory state is advantageous as compared to detection of heartbeat information representing a cardiac state. For example, the change in tension that occurs in the seat belt synchronizes with the movement of the chest and includes respiratory information. Therefore, by detecting a change in tension of the seat belt in a state in which the occupant wears the seat belt, it is possible to accurately detect breathing information without making the occupant feel bothersome.

Thus, the heart rate fluctuation is correlated with the respiration, and under the on-vehicle environment, the respiration information can be detected with high accuracy as compared to the heart rate information. Therefore, the estimation unit 30 of the present embodiment can obtain characteristic information related to the activity of the autonomic nerve of the occupant (for example, information corresponding to the above-described LF / HF obtained from the heart rate fluctuation) by the detection unit 40. It is extracted from the respiration information, and the awakening state of the occupant is estimated based on the characteristic information.

In addition, respiration changes also by movement (body movement) of the body whose displacement is larger than respiration. Therefore, if it is considered that body movement is also considered to be a part of respiratory activity, it may be considered that the state of body movement is added to the estimation of the arousal state of the occupant by using respiratory information to estimate the arousal state of the occupant. it can. Body movement includes, for example, posture change of the occupant on the seat.

Next, the configuration and functions of the present embodiment shown in FIG. 2 will be described in more detail.

The sensor 20 monitors a change in accordance with the breathing of the occupant 11 and outputs an output signal in accordance with the monitoring result. The change according to the breathing of the occupant 11 is, for example, a change in body movement synchronized with the breathing of the chest, belly, waist, back, back or buttocks, a change synchronized with the breathing such as the flow or temperature of the breath drawn from the nostril Etc. The detection unit 40 detects respiration information from a change according to the respiration of the occupant 11 acquired by the sensor 20. The sensor 20 is mounted on, for example, the seat 2, the seat belt 4, the buckle 8, the tongue 7 or a dashboard.

For example, the sensor 20 detects tension generated in the seat belt 4 (hereinafter, also referred to as “tension F”), and outputs an output signal that changes in accordance with the detected tension F. When the seat belt 4 is attached to the occupant 11, the movement of the chest or belly of the occupant 11 caused by the movement or breathing of the occupant 11 is transmitted to the seat belt 4, and the tension F of the seat belt 4 changes. The change in tension F of the seat belt 4 is transmitted to the tongue 7 and transmitted to the buckle 8 via the tongue 7. The sensor 20 may be provided on the main body 8 a of the buckle 8 or may be provided on the stay 8 b of the buckle 8. Thus, the sensor 20 may detect a change in accordance with the respiration of the occupant 11 as a change in tension T.

The sensor 20 detects, for example, a deformation or displacement caused by a change in the tension F of the seat belt 4 as the tension F of the seat belt 4. For example, the sensor 20 may be a strain sensor that detects a change in load input from the seat belt 4 to the buckle 8 via the tongue 7 or detects a change in capacitance generated by a change in tension F of the seat belt 4 It may be a capacitive sensor. When the tension F of the seat belt 4 changes, the buckle 8 itself connected to the seat belt 4 via the tongue 7 is displaced. Therefore, the sensor 20 may be a device that detects the displacement of the buckle 8 itself as a change in the tension F of the seat belt 4. For example, a non-contact sensor etc. which detect a relative distance with a reflective subject outside buckle 8 by sending and receiving of light or an electric wave are mentioned.

The chest movement of the occupant 11 mainly changes the tension of the shoulder belt portion 9, and the belly movement of the occupant 11 mainly changes the tension of the lap belt portion 10. Then, both the shoulder belt portion 9 and the lap belt portion 10 are connected to the buckle 8 via the tongue 7. Therefore, the sensor 20 provided on the buckle 8 can detect the information on the movement of both the chest and the belly of the occupant 11 from the change in tension, so that the sensor 20 is provided on the buckle 8 to improve the detection accuracy of the tension F. As a result, the detection accuracy of the breathing information of the occupant 11 is improved.

The detection unit 40 extracts a respiration signal including respiration information of the occupant 11 from the output signal of the sensor 20. For example, after checking whether the possible numerical range of the output signal of the sensor 20 is within the appropriate range, the detection unit 40 performs noise removal and filtering for selectively emphasizing the cycle and amplitude of the respiration signal. On the output signal of The period of steady breathing during driving is usually in the range of 3 seconds to 6 seconds, but the range differs depending on each person, and also depending on the awake state of each person or the like. Therefore, the detection unit 40 passes a signal of a frequency range (for example, a range from 0.04 Hz (25-second cycle) to 0.5 Hz (2-second cycle)) wider than the frequency range of steady breathing during driving, It is preferable to use a filter that selectively cuts signals of frequencies outside the frequency range.

FIG. 3 is a flowchart illustrating an example of the respiration signal extraction process performed by the detection unit 40. FIG. 4 is a flowchart showing an example of the respiratory cycle statistical process performed by the detection unit 40. FIG. 5 is a flowchart showing an example of the respiratory frequency component ratio detection process performed by the detection unit 40. The detection unit 40 repeatedly performs each of these processes shown in FIGS. 3 to 5 at a predetermined cycle. Next, each process shown in FIGS. 3 to 5 will be described.

FIG. 3 is a flowchart illustrating an example of the respiration signal extraction process performed by the detection unit 40. The detection unit 40 reads the output signal s output from the sensor 20 (step S11), and executes a process of extracting the respiration signal Rs from the read output signal s (step S13). The detection unit 40 passes signals in a frequency range from 0.04 Hz (25-second cycle) to 0.5 Hz (2-second cycle) using, for example, a low pass filter and a high pass filter, and signals with frequencies other than that range To the output signal s. Thereby, the respiration signal Rs is extracted from the output signal s.

In step S15, the detection unit 40 normalizes the extracted respiration signal Rs to generate a normalized respiration signal Rsn. The detection unit 40 normalizes the respiration signal Rs by performing a normalization process of adjusting the amplitude and the offset of the respiration signal Rs in order to increase the detection accuracy of the respiration cycle and the frequency component of the respiration signal Rs.

FIG. 6 is a diagram showing an example of the respiration signal Rs before normalization. FIG. 7 is a diagram showing an example of the respiration signal Rsn after normalization. For example, in order to increase the detection accuracy of the period and frequency components of the respiration signal Rs, the detection unit 40 sets the respiration signal Rs such that the amplitude center Rc of the respiration signal Rs becomes zero and the average amplitude Rsmave of the respiration signal Rs becomes one. Are normalized to generate a normalized respiration signal Rsn.

There are various methods for normalizing the amplitude of the respiration signal Rs depending on the purpose. When the purpose is to detect a respiratory cycle or a respiratory frequency component, the detection unit 40 may normalize the average amplitude Rsmave to 1, for example, in order to avoid the influence of the amplitude fluctuation, and the following simple Amplitude shaping and normalization may be performed by any method.

FIG. 8 is a diagram for explaining some simple methods of amplitude shaping normalization processing. FIG. 8A shows an example of the respiration signal Rs before normalization. FIG. 8 (b) shows an example of a normalization process for generating a normalized respiration signal Rs by limiting the amplitude of the respiration signal Rs to a predetermined level. FIG. 8C shows that by limiting the rate of change in amplitude of the respiration signal Rs at predetermined upper and lower levels, fast and slow changes in amplitude are suppressed and the amplitude is limited to a certain level. An example of the normalization process which produces | generates signal Rsn is shown. FIG. 8D shows an example of a normalization process for generating a breathing signal Rsn of an eye pattern by comparing the amplitude of the breathing signal Rs with the zero crossing and performing angle limitation on the changing edge. FIG. 8E shows an example of the original waveform of the respiration signal Rs when the amplitude of the respiration signal Rs is compared at the zero crossing.

FIG. 4 is a flowchart showing an example of the respiratory cycle statistical process performed by the detection unit 40. In step S21, the detection unit 40 detects a breathing cycle RI which is one of the breathing information from the normalized breathing signal Rsn, and generates a breathing cycle fluctuation RIV which is time series data of the breathing cycle RI. The detection unit 40 detects the breathing cycle RI by detecting the zero cross or peak of the normalized respiration signal Rsn.

In step S23, for example, the detection unit 40 performs statistical analysis of respiratory cycle fluctuation RIV in a predetermined section. The detection unit 40 detects, for example, respiratory information such as an average respiratory cycle RIsave, a respiratory cycle standard deviation RIsstd, an average difference fluctuation RIVsave, and an average amplitude Rsmave.

FIG. 9 is a diagram showing an example of statistical analysis of respiratory cycle fluctuation RIV in a predetermined section.

The detection unit 40 calculates an average value of measurement data S ₁ to S _n of n (n is an integer of 2 or more) respiratory cycles RI measured in a predetermined calculation interval as an average respiratory cycle RIsave.

The detection unit 40 calculates a standard deviation of measurement data S ₁ to S _n of _n respiratory cycles RI measured in a predetermined calculation interval as a respiratory cycle standard deviation RIsstd. The respiratory cycle standard deviation RIsstd represents the variation from the average value of the calculation interval, and decreases when the respiratory cycle RI is stable, and increases when the respiratory cycle RI changes. It can not be distinguished from the respiratory cycle standard deviation RIsstd whether the respiratory cycle RI fluctuates randomly or slowly in the calculation section.

The detection unit 40 calculates the square root of the squared integrated value of the difference change for each breath as the average difference fluctuation RIVsave. That is, “RIVsave = √ ((S ₂ −S ₁ ) ² + (S ₃ −S ₂ ) ² +... (S _n −S _n−1 ) ² )”. The average difference fluctuation RIVsave increases when the breathing cycle RI fluctuates randomly in the calculation section, and decreases when the breathing cycle RI fluctuates slowly and greatly in the calculation section.

The detection unit 40 calculates an average value of the amplitudes of the respiration signal as an average amplitude Rsmave. For example, when the amplitude of the respiration signal Rs is twice or more the average amplitude Rsmave, the determination unit 70 determines that the respiration is a respiration state different from a normal state such as deep respiration or body movement.

FIG. 5 is a flowchart showing an example of the respiratory frequency component ratio detection process performed by the detection unit 40. When the respiratory frequency component ratio detection process is performed, the respiratory signal Rs does not necessarily have to be normalized, and the respiratory signal Rs may be used as it is in the respiratory frequency component ratio detection process. However, in this embodiment, the frequency analysis is performed using the respiration signal Rsn normalized in step S15 of FIG. 3 to reduce the influence of the amplitude fluctuation.

In step S31, the detection unit 40 samples the respiration signal Rs (or the respiration signal Rsn after normalization) at a predetermined sampling frequency fs. For example, the detection unit 40 samples the respiration signal Rs (or the respiration signal Rsn after normalization) at the sampling frequency fs (= 4 Hz) to create 2 ⁸ (= 256) or more pieces of time-series data. If the sampling frequency fs is 10 Hz, it is preferable to create 2 ⁹ (= 512) or more pieces of time-series data (for example, 2 ¹⁰ (= 1024)).

In step S33, the detection unit 40 analyzes frequency components of the respiration signal Rs in the range of 0.04 Hz to 0.5 Hz. The detection unit 40 performs fast Fourier transform (FFT) by multiplying time series data of the number of powers of 2 by the same number of window functions. The detection unit 40 performs an FFT on the respiration signal Rs to integrate the integrated value LFR of the power spectrum amplitude of the low frequency respiration component of the respiration signal Rs and the integrated value HFR of the power spectrum amplitude of the high frequency respiration component of the respiration signal Rs. And calculate. The detection unit 40 preferably calculates, for example, an integrated value LFR of power spectrum amplitudes of low frequency respiration components in the same frequency range (range from 0.04 Hz to 0.15 Hz) as the above-described LF. In addition, it is preferable that the detection unit 40 calculate, for example, an integrated value HFR of power spectrum amplitudes of high frequency breathing components in the same frequency range (range of 0.15 Hz to 0.5 Hz) as the above-described HF. The power spectrum is calculated as a real number, for example, after FFT calculation of the respiration signal Rsn, multiplied by a complex conjugate.

When the sampling frequency fs is 4 Hz and the time series data of 2 ⁸ (= 256) or more, the minimum frequency resolution is 0.033 Hz (= 2 × 4/256), which is a value smaller than 0.04 Hz, The number of time series data is appropriate. In the case of sampling frequency fs of 4 Hz and more than 2 ⁹ (= 512) time series data, the minimum frequency resolution is 0.016 Hz (= 2 × 4/512), so the value is smaller than 0.04 Hz. The number of time series data is more appropriate.

In step S35, the detection unit 40 divides the LFR by the HFR to calculate the respiratory frequency component ratio RLHR, which is the ratio of LFR to HFR (LFR / HFR). The detection unit 40 performs a filtering process using an infinite impulse response filter on the RLHR, calculates a time average RLHRave of the past RLHR, and calculates RLHRn (= RLHR / RLHRave). RLHRn is a normalized output of the respiratory frequency component ratio.

FIG. 10 is a waveform diagram showing an example of each signal detected from the driver who is driving. FIG. 10 (a) is a waveform diagram showing the respiration signal Rs and the heartbeat interval RRI. The heart rate interval RRI is a value measured by an electrocardiograph. The horizontal axis represents data points, and the vertical axis of RRI represents heart rate per minute. As shown in FIG. 10A, the change in respiration and the heart rate fluctuation are linked.

FIG. 10 (b) is a waveform diagram showing LF / HF and LFR / HFR. LF / HF indicates the value of LF / HF obtained from the heart rate interval RRI measured by the electrocardiograph, and is normalized so that the average value of all the sections becomes 1. On the other hand, LFR / HFR indicates the value of LFR / HFR obtained from the respiration signal Rs in the above-described respiration frequency component ratio detection processing of FIG. 5 and is normalized so that the average value over the entire interval is 1. . As shown in FIG. 10 (b), LFR / HFR shifts similarly to LF / HF.

FIG. 10 (c) shows a breathing cycle RI. The unit of the vertical axis is seconds.

As shown in FIG. 10, RLHR (= LFR / HFR) or RLHRn (normalized output of respiratory frequency component ratio) calculated by respiratory frequency component ratio detection processing of FIG. The change is similar to the value of LF / HF calculated based on the interval RRI. That is, RLHR (= LFR / HFR) or RLHRn (normalized output of respiratory frequency component ratio) calculated based on the respiration signal Rs is similar to the sympathetic nervous system and the secondary sympathy, as LF / HF calculated based on the heart rate interval RRI. It can be used as index data that indicates neural activity.

Therefore, instead of LF / HF, RLHR or RLHRn can be obtained by making the range of the frequency ratio of RLHR or RLHRn calculated from the respiration signal Rs the same as the range of the frequency ratio calculated from the heart rate interval RRI. Can be used to estimate arousal.

In particular, since respiration is closely related to parasympathetic activity, in a state where sympathetic activity is reduced and parasympathetic activity is activated, the autonomic nervous activity estimated from respiratory signal Rs is It is considered to be in good agreement with the autonomic nerve activity estimated from the interval RRI. This is because the parasympathetic nerve predominates in a period in which the value of RLHR or RLHRn is relatively low, and the correlation with respiration is higher. Therefore, for example, when the value of RLHR or RLHRn is lower than a predetermined determination threshold, the determination unit 70 can determine that the autonomic nerve is in the parasympathetic dominant state and the arousal state is reduced. For example, in FIG. 10B, when the value of RLHRn is lower than the predetermined determination threshold value 0.4, the determination unit 70 determines that the autonomic nerve is in the parasympathetic dominant state, and the awakening state decreases. It is determined that there is.

In addition, during a period in which the value of RLHR or RLHRn is relatively low, arousal may be reduced and sleepiness may occur after the relaxation state is reached. When the drowsiness increases and the occupant notices sleepiness, the occupant begins to make efforts to wake up (awakening effort), the arousing sympathetic nerve becomes active, and the activity balance with the parasympathetic nerve largely fluctuates. Therefore, the determination unit 70 can determine that drowsiness has occurred in the occupant when the repeated pattern that increases after the value of RLHR or RLHRn decreases below the predetermined determination threshold is obtained from the detection unit 40.

Further, RLHR (= LFR / HFR) shown in FIG. 10 (b) represents RLHRn normalized so that the average value of all the displayed sections is 1. Converting the frequency range for calculating RLHRn to a respiration cycle, HFR corresponds to a respiration cycle of 2 to 6 seconds, and LFR corresponds to a respiration cycle of 6 to 25 seconds. HFR can be understood as a cycle of respiration itself, and LFR can be understood as a periodic increase / decrease fluctuation component of a respiration cycle. Therefore, the value of RLHR (= LFR / HFR) fluctuates due to a change in the ratio of LFR to HFR even when the average respiratory cycle changes in size, and the absolute value of RLFR tends to increase as the respiratory cycle is delayed. Therefore, normalization is preferable to reliably capture changes in respiration.

For example, normalization is performed by correlating the average period of respiration and the magnitude of RLHR and dividing RLHR by the average period according to the correlation coefficient. When focusing on changes in RLHR, normalization can be performed by dividing RLHR by the average value of the past several minutes of the RLHR or the infinite impulse response filter value. When normalized in this manner, changes in the average respiratory cycle of each person can be absorbed, and when there is a change in RLHR, the change appears emphasized in RLHRn. Therefore, the estimation accuracy can be enhanced by using RLHRn to estimate the awake state.

Also, in general, deep and slow breathing is interpreted as relaxing and low arousal. As one example, when changing from 5-second cycle breathing to 8-second cycle breathing, it indicates that the breathing becomes slow, so this usually corresponds to a decrease in arousal. The 5-second period coincides with the frequency domain of HFR, and the 8-second period coincides with the frequency domain of LFR. In this case, the change from 5-second cycle breathing to 8-second cycle respiration may be a transition from a large HFR state to a large LFR state, and a transition from a small LFR / HFR state to a large state. That is, according to the LFR / HFR values, the arousal level is increased, so it seems to be in contradiction to the arousal decrease.

However, in general, if the breathing is fast, the inspiratory time and the exhaust time are almost the same, but if the breathing is longer, the exhaust time is longer. That is, the intake time does not change much. Then, assuming that the time taken for intake is 2 seconds to 3 seconds, the remaining time is exhausted, so the frequency component by intake increases the power spectrum in the HFR region. Since the repetition cycle of the inspiration is added as a power spectrum of LFR, both the denominator and the numerator of RLHR (= LFR / HFR) are simultaneously increased, resulting in balance. After all, even if breathing is performed for a long period, the LFR corresponding to the long period is not only increased but also the HFR is increased. As a result, it is less likely that the arousal level is greatly increased. Therefore, in the present embodiment, the awake state can be determined with high accuracy by analyzing the frequency components of respiration and determining the awake state, as compared to the method of determining the awake state by focusing only on the respiration cycle. With such a mechanism, the activity of autonomic nerve is associated with respiration.

In the above-mentioned embodiment, although the method of using the FFT to analyze the frequency component of respiration is used, it is possible to use the simple low pass filter, high pass filter and slide window filter without using the FFT. Frequency components can be analyzed. The simplest discrete low-pass and high-pass filters are infinite impulse response filters, and the characteristics of an analog CR filter can be realized by simple calculation of difference equations. The frequency range of LFR and HFR is as shown in FIGS. 11 and 12 in combination of the low pass filter and the high pass filter. However, since the filter cutoff characteristics are not steep, the contrast between LFR and HFR is reduced. Therefore, the filter characteristics can be improved by superposing a filter having a notch characteristic in the vicinity of 0.15 Hz.

A convolution filter can be considered as a filter having a notch characteristic which is easy to calculate. Next, the convolution filter will be described with reference to FIGS.

The data sequence of the respiration signal Rsn is S0, S-1, S-2, ... S-n. The detection unit 40 uses a rectangular function as the simplest embodiment, and translates the data sequence into S0, S-1, S-2,. Perform a convolution operation that adds and adds to. In the example, the origin is T = 0. Since the amplitude of the rectangular function is 1 or −1, the detection unit 40 may divide the sum of Sn in the section in which the rectangular function is 1 by the number of data in the section. In a section where the amplitude of the rectangular function is −1, the detection unit 40 multiplies the data of the section by −1 and calculates the sum of Sn. The filter characteristics are determined by the sampling frequency and the numerical selection of n0, n1, n2, n3, n4 and n5 (for example, sampling frequency = 4 Hz, n0 = 0, n1 = 13, n2 = 4, n3 = 10, n4 = 19) , N5 = 39). FIG. 14 is a schematic view of F2-F1 filter characteristics and F3 filter characteristics. FIG. 15 is a diagram showing superposition of a convolution filter, a low pass filter and a high pass filter.

FIG. 16 is a diagram showing an example of a result of calculating RLHRn (normalized output of respiratory frequency component ratio) using the filter of FIG. FIG. 16 is the same as FIG. 10 except that the waveform of RLHRn is added to FIG. Compared with the calculation result by the method of analyzing the frequency component of respiration using FFT, the contrast is lowered because the frequency discrimination ability of the filter is rough, but the tendency of increase or decrease of RLHRn almost agrees with LF / HF. ing. Therefore, it is possible to estimate the activity of the autonomic nerve even if a filter having a notch characteristic which is easy to calculate is used.

Next, an embodiment in which the amount of calculation of the respiratory frequency component ratio is suppressed will be described. This embodiment is suitable for implementation in a computing environment such as an 8-bit micro processing unit (MPU) or the like in which memory or computation accuracy is insufficient.

As illustrated in FIG. 8E, the detection unit 40 converts the amplitude of the respiration signal Rs into a binary signal of 1 and 10, and calculates one respiration cycle for each rising edge or falling edge of the binary signal. . Alternatively, the detection unit 40 calculates one breathing cycle at each change point (half cycle) of the binary signal. FIG. 17 shows time-series data of 20 respiratory cycles obtained in this manner (N = 20). FIG. 18 shows a frequency distribution (frequency of occurrence) in which the horizontal axis is a cycle interval of 0.5 second intervals for the 20 respiratory cycle data. Since the respiratory cycle directly corresponds to the fundamental frequency of respiration, the reciprocal of the cycle corresponds to the fundamental frequency of respiration. The frequency is 0.25 Hz for 4 seconds and 0.166 Hz for 6 seconds. FIG. 19 shows a frequency distribution (occurrence frequency) in which the horizontal axis of the frequency distribution of FIG. 18 is sectioned by frequency and rearranged. The frequency distribution of FIG. 19 matches the FFT analysis result of the fundamental frequency of respiration.

Therefore, it is possible to simply analyze the frequency and estimate the ratio of respiratory frequency components using the frequency distribution of the respiratory cycle. In the example of the figure, when the sum of frequencies from 0.04 Hz to 0.15 Hz is LF_bin and the sum of frequencies from 0.16 Hz to 0.5 Hz is HF_bin, LF_bin = 1 and HF_bin = 19, respectively. Therefore, LFR / HFR = 1/19 = 0.05. This is a simple indicator of autonomic nerve activity determined from the respiratory cycle. However, if the respiratory cycle is 6 seconds or less and stable, then LF_bin = 0 and LF / HF = 0. Further, in the case where the breathing cycle is stable for 7 seconds or more, LF_bin = 20, and it becomes an unnatural determination such as being always active at LF / HF = 20.

FIG. 20 shows an example of a breathing waveform with a repetition cycle of 6 seconds. FIG. 21 shows an example of the result of spectral analysis of the waveform of FIG. In FIG. 21, the sampling frequency is 2 Hz, and FFT analysis is performed on data of 64 points. Two spectral peaks are observed around 0.16 Hz and 0.33 Hz. 0.16 Hz corresponds to a repetition cycle of 6 seconds of the respiration waveform, which is a fundamental frequency of one respiration. 0.33 Hz corresponds to a harmonic spectrum associated with imbalanced time of intake and exhaust time, and is a second harmonic component with a period of 3 seconds. This second harmonic component corresponds to the occurrence of an imbalance in which the duty ratio of the waveform after shaping in FIG. 20 does not become 50%. In particular, the longer the cycle, the greater the imbalance.

The simplified frequency ratio calculation method using the frequency distribution estimates the frequency component of the respiration amplitude from the time series of the respiration cycle based on a predetermined rule. Then, the frequency component of one respiration waveform is distributed to the HFR component, the frequency component of time series fluctuation of respiration is distributed to the LFR component, the frequency distribution is generated, and the respiratory frequency component ratio is estimated by the ratio of LFR and HFR frequency. Do.

Next, an example of the rule of the simple estimation of the respiratory frequency component ratio will be described with reference to FIGS. The periodic sequence is I (0) to I (-20) (the number n in parentheses is the n-th breath from the present to the past, and I (n) is its respiratory cycle).

(0) The detection unit 40 reads respiration amplitude data from the respiration signal Rs (step S41). The detection unit 40 repeatedly calculates one respiratory cycle based on the change in the read respiratory amplitude data, and creates a signal sequence of 21 respiratory cycles from I (0) to I (-20) (step S43). ).

(1) The detection unit 40 counts respiration cycles of 2 seconds to 6 seconds as frequencies of HF_bin, including harmonic spectra (from step S45 to step S51).

(2) On the other hand, the breathing cycle of 6 to 12 seconds has a fundamental frequency of 6 seconds or more and a harmonic frequency of 6 seconds or less because the second harmonic and third harmonic components are 6 seconds or less. . Therefore, the detection unit 40 counts a breathing cycle of 6 to 12 seconds as the frequency of both LF_bin and HF_bin (from step S49 to step S55). The ratio of incorporation into LF_bin and HF_bin (that is, the ratio of LF_bin and HF_bin as a frequency to be incorporated) may be determined based on the duty ratio of one cycle. In the example, the incorporation ratio is set to 1: 1 for the sake of simplicity. However, since the calculation of the frequency integration changes and a discontinuity occurs at the boundary of 6 seconds, in the cycle of 5 seconds to 7 seconds (Step S49 Yes), the integration balance to LF_bin and HF_bin is continuously changed according to the cycle By this (step S51), the discontinuity is reduced.

That is, in step S45, the detection unit 40 determines whether I (n) is data of a breathing cycle less than 5 seconds. If the detection unit 40 determines that I (n) is data of a respiration cycle less than 5 seconds (Yes at step S45), the detection unit 40 increments HF_bin by one (step S47). On the other hand, when the detecting unit 40 determines that I (n) is data of a breathing cycle of 5 seconds or more (No in step S45), the detecting unit 40 does not increment HF_bin. Next, in step S49, the detection unit 40 determines whether I (n) is data of a breathing cycle of 5 seconds or more and less than 7 seconds. When the detecting unit 40 determines that I (n) is data of a breathing cycle of 5 seconds or more and less than 7 seconds (Yes in step S49), the detecting unit 40 calculates HF_bin and LF_bin according to the two equations described in step S51. Next, the detection unit 40 determines whether I (n) is data of a breathing cycle of 7 seconds to 12 seconds (step S53). If the detecting unit 40 determines that I (n) is data of a breathing cycle of 7 seconds or more and 12 seconds or less (Yes at step S53), it increments LF_bin by one (step S55). On the other hand, when the detection unit 40 determines that I (n) is not data of a breathing cycle of 7 seconds or more and 12 seconds or less (No in step S53), the detection unit 40 does not increment LF_bin. The detection unit 40 performs the processing from step S45 to step S55 for each of the 21 respiratory cycle data from I (0) to I (-20).

(3) When the sum A of the two breathing cycles I (n) and I (n-1) is 6 seconds or more and 12 seconds or less, the detection unit 40 calculates the difference B (= abs (I (n) -I). (n-1))) is evaluated (from step S61 to step S70). "Abs (*)" represents the absolute value of *. That is, when the difference B is smaller than one second (Yes at step S69), the detection unit 40 determines that I (n) and I (n-1) have frequency components with the same cycle of six seconds or less, and the frequency of HF_bin. Is incremented and counted (step S70). If the difference B is larger than 2 seconds (Yes at step S65), the detecting unit 40 determines that I (n) and I (n-1) have frequency components of 6 seconds to 12 seconds, and the LF bin frequency is one And increment (step S67).

(4) The detection unit 40 calculates the sum A (= I (n) of the two breathing cycles when the sum A of the two breathing cycles I (n) and I (n-1) is not less than 6 seconds and not more than 24 seconds. Temporarily consider + I (n-1) as one absorption cycle, and the difference C (= (I (n) + I (n-1))-(I (n-2) + I (n-3)) Is evaluated (from step S71 to step S79). That is, when the difference C is smaller than one second (Yes at step S77), the detection unit 40 determines four respiratory cycles (I (n), I (n-1), I (n-2), I (n-). 3) It is judged that there is no long cycle fluctuation, and those four breathing cycles have frequency components of similar cycles of 6 seconds or less. In this case, the detection unit 40 increments the frequency of HF_bin by one and counts (step S79). If the difference C is greater than 3 seconds (Yes at step S73), the detection unit 40 selects four respiratory cycles (I (n), I (n-1), I (n-2), I (n-3). ), And it is determined that the four breathing cycles have frequency components of 6 seconds or more and 24 seconds or less. In this case, the detection unit 40 increments the frequency of LF_bin by one and counts it (step S75).

(5) The detecting unit 40 performs the processing from step S61 to step S79 for each of n from 0 to -17 (step S81). Then, the detection unit 40 calculates LF / HF_bin corresponding to LFR / HFR by dividing LF_bin by HF_bin (step S83). In addition, (4) can be extended similarly to three or more and calculated similarly also with respect to the sum of three breathing cycles. In addition, in the incorporation into LF_bin and HF_bin, weighting factors are added and incorporated respectively in (1) to (5), but in this example, all were calculated as weighting factor 1.

FIG. 24 is a diagram showing an example of a result of calculating RLHRn (normalized output of respiratory frequency component ratio) using a simple estimation method of respiratory frequency component ratio. FIG. 24 is the same as FIG. 16 except that the waveform of RLHRn according to the simple estimation method is added to FIG. 24 (b). The waveform of RLHRn based on the simple estimation method is time series data of 10 respiratory cycles, and represents the result calculated for each determination of one respiratory cycle. Compared with the calculation result by the method of analyzing the frequency component of respiration using FFT, the tendency of increase and decrease of RLHRn by the simple estimation method is almost in agreement with LF / HF. Therefore, it is possible to estimate the activity of the autonomic nerve even by using a simple estimation method of the respiratory frequency component ratio.

The method of estimating the respiratory frequency component from the respiratory signal is not limited to the above. For example, the threshold value for dividing LFR and HFR is 6 seconds (0.16 Hz) in order to correspond to the activity index of the autonomic nerve estimated from the heart rate interval RRI. For example, even if the threshold is 4 seconds or 8 seconds, detection of the feature is possible. The division range of the frequency frequency may be increased. Further, if the respiratory rate to be analyzed is constant, LFR may be used as a value corresponding to LF / HF without calculating the ratio (RLHR) of LFR to HFR. That is, if LFR can be used as index data indicating the activity of the sympathetic nerve and the parasympathetic nerve, it may be used as an index used to determine the arousal state. Thus, the index used to determine the awake state is not limited to the ratio of LFR to HFR (RLHR).

The threshold for estimating frequency components of fluctuations over a plurality of cycles, such as 2 cycles or 3 cycles, and the difference calculation method can be similarly changed in accordance with the purpose of feature detection. Moreover, in order to suppress the discontinuous change of the numerical value due to the change of the judgment condition before and after the judgment threshold value of fluctuation over 2 cycles or 3 cycles, it is possible to apply the method of continuously changing the balance illustrated in the case of 1 cycle is there. Three examples of the FFT approach, the filter approach and the frequency distribution approach are described above. In either method, the frequency component of one amplitude cycle of the respiration signal and the frequency component of the time-series repeated fluctuation of respiration are analyzed, the numerical value characterizing respiration is taken out from the analysis result, the estimation of the respiration state, Estimate autonomic nervous activity from respiration. The respiratory state and the nerve activity state taken out by these methods are useful for estimating the arousal state and stress state of the occupant.

FIG. 25 is a block diagram showing an example of the configuration of a buckle in the second embodiment. The description of the configuration and effects similar to those of the first embodiment in the second embodiment will be omitted or simplified by using the above description. In the first embodiment (see FIG. 2), the determination unit 70 and the output unit 80 are provided in the buckle 8, whereas in the second embodiment (see FIG. 25), the determination unit 70 and the output unit The reference numeral 80 is provided on a device different from the buckle 8. The awake state determination system 15 shown in FIG. 25 includes a buckle 8 and an estimation unit 30. The estimation unit 30 includes a detection unit 40, a determination unit 70, and an output unit 80.

As mentioned above, although the buckle, the awakening state determination system, and the awakening state determination method were demonstrated by embodiment, this invention is not limited to the said embodiment. Various modifications and improvements, such as combinations or permutations with part or all of the other embodiments, are possible within the scope of the present invention.

For example, the seat 2 may be a front seat or a rear seat of a vehicle. Further, in FIG. 25, the detection unit 40 may be provided in a device other than the buckle 8. In addition, a part of the plurality of processes performed on the output signal s of the sensor 20 (for example, the normalization process) may be performed by the determination unit 70 instead of the detection unit 40.

This international application claims priority based on Japanese Patent Application No. 2018-006304 filed on Jan. 18, 2018, and the entire contents of Japanese Patent Application No. 2018-006304 In the

Reference Signs List 1 seat belt device 8 buckle 15 awake state determination system 20 sensor 30 estimation unit 40 detection unit 70 determination unit 80 output unit

Claims

A sensor that outputs an output signal that changes in response to the breathing of a vehicle occupant;
A detection unit that detects frequency components of the respiration from the output signal by frequency analysis;
A buckle comprising: a determination unit that determines an awake state of the occupant based on the frequency component detected by the detection unit.
The detection unit uses the frequency component to calculate index data indicating activity of sympathetic nerve and parasympathetic nerve.
The buckle according to claim 1, wherein the determination unit determines the awake state of the occupant based on the index data calculated by the detection unit.
The buckle according to claim 2, wherein the determination unit determines that the awakening state of the occupant is reduced when the index data is lower than a predetermined determination threshold.
The determination unit according to claim 2, wherein the drowsiness is generated in the occupant, when the repetitive pattern which rises after the index data decreases below a predetermined determination threshold is obtained from the detection unit. buckle.
The index data is a ratio of an integrated value of an amplitude of a power spectrum of the low frequency component of the respiration to an integrated value of an amplitude of a power spectrum of the high frequency component of the respiration whose frequency range is higher than the low frequency component. A buckle according to claim 2.
The buckle according to claim 2, wherein the index data is normalized data.
A buckle having a sensor that outputs an output signal that changes in response to the breathing of a vehicle occupant;
A detection unit that detects frequency components of the respiration from the output signal by frequency analysis;
And a determination unit that determines the awake state of the occupant based on the frequency component detected by the detection unit.
A sensor provided on the buckle outputs an output signal that changes in response to the breathing of the vehicle occupant,
The detection unit detects frequency components of the respiration from the output signal by frequency analysis;
The awake state determination method, wherein the determination unit determines the awake state of the occupant based on the frequency component detected by the detection unit.